Current nuclear medicine provides various techniques for non-invasive diagnosis of internal physical structures and biochemical processes occurring within a patient. Computerized axial tomography (CAT), magnetic resonance imaging (MRI) scans and conventional X-ray methods are examples of such structure-imaging systems. These techniques, which identify and localize only physical structures, suffer from the drawback that by the time an abnormality appears, and is detected, the pathological condition creating such an abnormality is often well advanced.
In contrast, positron-emission tomography (PET) systems are used to image functioning metabolic systems in the brain as well as the rest of the body. By imaging function rather than structure, these systems provide a unique complement to X-ray, CAT and MRI systems. PET is accomplished through the coincident detection of pairs of gamma rays. These gamma rays are produced when positrons emitted from the source (which is typically placed within the patient's body) annihilate with electrons in the tissues surrounding the source location. When the gamma rays are emitted at annihilation, the physical properties of these subatomic particles dictate that the two gamma rays are emitted at a single energy and with other known properties; for example, it is known that the gamma rays will be emitted such that they travel in directions very nearly opposite one another.
The PET imaging process includes a number of steps. Proton rich radioisotopes are first placed within the patient's body by, for example, injection or ingestion. In most cases these isotopes will be localized at or near the area where diagnosis is desired. Once incorporated into the body, the isotope continues to emit positrons as part of a naturally occurring decay process. The positron is an antielectron that, after traveling a short distance, will combine with an electron from the surrounding tissue and annihilate. On annihilation, the masses of both the electron and positron will be converted into electromagnetic radiation. In order to conserve energy and linear momentum, the electromagnetic radiation is in the form of two gamma rays which are of equal energy and which are emitted approximately 180 degrees to each other. It is this annihilation radiation that is detected externally in a PET device in order to measure both the quantity and the location of the positron emitter as it moves through the body.
The concentration of the radioisotope as it moves through and is processed by the patient's body can be measured and displayed as a cross section gray scale image. In this image, the intensity of each pixel (picture element) is proportional to the concentration of the radioisotope at that position within the body. This type of so called "kinematic" technique has been and will likely continue to be one of the most powerful methods for diagnosing and analyzing dynamic processes such as blood flow, substrate transport and biochemical reactions within the human body.
PET systems currently existing can record and process a large number of tomographic images of a human brain or torso simultaneously. Moreover, sensors can be placed either in a planar ring structure capable of forming a two dimensional image or in a volumetric layout to achieve a three dimensional image. The latter layout is termed positron volume imaging (PVI), although some authors will use the term "PET" when referring to PVI as well. PVI can be set up to process data initially as PET data, later combining planar images to form a volumetric image. Alternatively, PVI can be achieved by permitting inter-plane coincidences at the sensors and processing these coincident detections accordingly. The obtainable resolution with either of these systems has been recently narrowed to under one centimeter, and the distribution of radioactivity within the subject can be assessed to within a few percent.
In a typical PET implantation, a ring of gamma ray sensors are positioned to surround the patient in a position local to the radioisotope source. The detection process takes advantage of both the fact that gamma ray emission occurs at 180 degrees to each other and the fact that gamma rays are created simultaneously. Simultaneous or coincidence detection of the gamma ray by sensors on opposite sides of the patient places the site of the annihilation on or near a line connecting the centers of the two sensors. If only one detection takes place, then the annihilation has typically occurred outside of the volume or plane between the two detectors. In this case no event is recorded, since the source would be located outside of the diagnosed area.
An operational PET system typically includes the above described data acquisition subsystem including the radiation sensors and their associated circuitry, a fast computer with the necessary imaging software, and large amounts of memory for storing and processing sensor and other input data. A display system for immediate viewing of the image is also typically provided. Finally, a means for interactive processing and system control by the user is generally included.
As early as 1986, dozens of regional cyclotron-PET centers were in operation or under development worldwide, and that number continues to grow. A cyclotron-PET center typically consists of an accelerator (usually a small medical cyclotron) for generating radioisotopes, a positron emission tomograph (PET) and a chemistry laboratory for the synthesis of short-lived biological radiotracers. In the U.S., many such centers may be found at university-based medical research centers. The cost of a modern high resolution PET detector is more than $1 million, including approximately $100,000 for crystals and approximately $250,000 for photosensors (typically photomultipliers).
Since PET was first implemented in the 1970 's, it has undergone successive refinements. Unfortunately, the newest high-resolution PET systems operate at or near the intrinsic limitations imposed by the physics of this technology. Statistical limits from limited patient exposure to positron-emitting radioisotopes and source position-smearing from positron range and residual momentum at annihilation serve to limit image resolution and accuracy. In addition, systems operating near this intrinsic limit are quite complex and very expensive. This is due, in part, to the requirement for a very large number of sensor elements to achieve the desired resolution.
Current PET detection techniques also suffer from various inaccuracies that result from decreases in crystal width as higher resolutions are sought. These inaccuracies are generally termed "imaging artifacts." One such artifact is that of radial blurring, which results from crystal penetration from sources away from the axis of the system. In other words, if the line of coincidence is located at some distance from the diameter of the detector ring, the gamma ray may pass through one or several crystals before being absorbed by the detecting crystal. This, in turn, causes a broadening of the coincidence aperture function towards the edges of the field of view. This problem is additionally complicated as the attenuation length of the crystal material increases. In order to achieve equivalent efficiency when using a crystal material having a longer attenuation length (where such crystal may have otherwise desirable properties such as high brightness, high speed, or low cost) the corresponding crystal depth must be increased. But once the crystal is deepened, radial blurring is increased. Even with crystals with the shortest attenuation lengths in current use, radial blurring limits system resolution for objects a few centimeters from the central axis of the detector. There have been various proposals made in an attempt to solve this problem, with the primary solution being the use of a depth of interaction measurement for the photons interacting within the detector.
Modern imaging systems have attempted to minimize imaging artifacts by using dense scintillation crystals such as bismuth germinate (BGO), by employing very narrow crystals, and by using specialized sensors to determine the particular location of interaction of the gamma ray in larger crystals. A variety of methods have been proposed to accomplish depth of interaction measurements in very high resolution PET detectors, but such measurements all have required either many additional photosensors (such as photodiodes) with their associated electronics, or complex coding schemes.
Another class of imaging instruments used in clinical nuclear medicine applications is that of single photon emission computed tomography (SPECT) systems. A detector used for SPECT can have many attributes in common with a PET detector, although position resolution requirements are typically much less demanding. In the SPECT imaging process, a radioactive tracer is first placed within the patient's body by injection or ingestion. This radioisotope decays by continually emitting low energy gamma rays (photons) as it travels throughout the patient's body. It is this photon radiation that is detected externally by the SPECT device. The photon radiation energy detected by SPECT devices is typically between the range of 55-400 keV, which is lower than the annihilation gamma's (511 keV) in PET systems. The most widely used radionuclide is an isomer of technetium, .sub.m.sup.99 Tc, which has a half-life of 6 hours (the time required for exactly half of the radionuclide initially present to decay). The radionuclide decays by continually emitting gamma rays, in the case of .sub.m.sup.99 Tc, the gamma energy being 140 keV. Some of the commonly used isotopes are listed here with the decay photon energy and corresponding bodily imaging function:
.sup.201 T1, 80 keV; used for heart and tumor imaging. PA1 .sup.176 Ta, 55-65 keV, used for imaging the heart. PA1 .sup.133 Xe, 80 keV, used for lung and ventilation studies.
Conventional SPECT implemention calls for the use of a collimator, usually consisting of a thick lead sheet perforated with thousands of small holes, placed directly in front of a gamma camera (crystal detector). Generally, the collimator holes are perpendicular to the crystal so as to block the passage of obliquely incident photons to the crystal detector and to thereby select the direction of the incident photon. By rotating the gamma ray camera and/or the collimator around the patient, a series of two-dimensional projections can be formed from different directions. By applying various reconstruction techniques, the internal distribution of radioactive tracers can be recovered simultaneously for parallel two-dimensional transverse sections. This SPECT technique can then be used for three-dimensional imaging of radioactive tracer distributions located in the lungs, heart and brain. Conventional gamma cameras for use in SPECT generally employ from 36 to 90 photomultipliers as photosensors, with their associated readout electronics. In part due to this internal complexity, commercial SPECT systems range in cost from approximately $200,000 to $500,000.